Signal profiling for medical imaging systems

ABSTRACT

Scatter effects are reduced in a radiographic imaging device, such as a digital slot scan mammographic imaging device, by reducing detected scatter and processing detector information to compensate for scatter effects. In one embodiment, a digital mammographic imaging system ( 10 ) includes a source ( 24 ) for transmitting a narrow beam ( 28 ) and a detector assembly ( 32 ) for detecting the beam ( 28 ). The beam ( 28 ) and the detector assembly ( 32 ) are synchronously scanned across the patient&#39;s breast ( 48 ) to obtain an image. Collimator slats ( 74 ) are provided at the leading and trailing edges of the detector to reduce detected scatter. Additionally, attenuators ( 76  and  92 ) are provided at the ends of the scanned motion and at the anterior edge of the detector array to assist in determining a spatial intensity profile. The spatial intensity profile information together with other imaging signal and patient dependent parameters are used in image processing to estimate and compensate for various scatter effects including single and multiple scatters and Compton and Rayleigh scatter.

FIELD OF THE INVENTION

[0001] The present invention generally relates to medical imagingsystems, and more particularly, to a method and system for enhancingradiographic images by removing or otherwise reducing the effects ofscatter.

BACKGROUND OF THE INVENTION

[0002] Medical imaging is considered to be one of the most demandingapplications of radiographic imaging, at least in terms of resolutionand/or contrast. For instance, in the case of mammography it isgenerally desirable that a given imaging system be capable of obtainingimages with a resolution on the order of fifty microns or better.Contrast requirements are also quite demanding since potential lesionsand/or suspicious masses may exhibit x-ray attenuation characteristicssimilar to that of surrounding “healthy” tissue. With early detection ofthese lesions and/or masses being extremely desirable, enhancement ofimage resolution and/or contrast continues to be of increasingimportance.

[0003] Mammography has been performed using both film-based and digitalsystems. In film-based systems, x-rays signals are generally transmittedthrough mammary tissue and received at an appropriate screen (e.g.,phosphor screen). Light emitted from the screen, due to excitation bythe impinging x-rays, is used to expose a light sensitive film. The filmis then developed to yield an image of the patient's breast which can beviewed on a light box. By contrast, a full field light-sensitivedetector may be utilized in place of the film in digital systems. Inthis context, “full field” indicates that the field imaged correspondsto the dimensions of the detector, although this may be substantiallyless than a full image area of interest to a physician, such as a fullbreast in the case of mammography. This detector outputs an electronicsignal that is indicative of the received radiation intensity across apixel region. In turn, the signal may be connected and processed into aviewable digital image (e.g., on a high-resolution monitor). As may beappreciated, digital image systems are becoming the norm in view of theattendant image storage and processing advantages.

[0004] Regardless of the mammography system utilized, granularity andscreen noise may tend limit and/or distort the resultant image. Forinstance, the image contrast has been shown to be significantly affectedby scattered radiation. Indeed, as a radiation source transmits a beamtowards a tissue, the beam is both attenuated and scattered by thetissue. The scattered radiation that impinges on the detector (e.g. inthe case of a digital system) from a path outside a “direct” orsubstantially straight path from the radiation source to the detector(“primary ray”) is generally undesirable. Accordingly, mammographicimages would ideally be generated “scatter-free”.

[0005] One type of scatter is Compton scattering, also referred to asincoherent scattering, that typically occurs when an incident x-rayphoton ejects an electron from an atom (e.g., a tissular atom), and anx-ray photon of lower energy is “scattered” from the atom. It may besaid that relativistic energy and momentum are conserved during Comptonscattering and that the scattered x-ray photon generally has less energyand therefore greater wavelength than the original incident photon. Bycontrast, Rayleigh scattering, also referred to as Thomson orcoherent/classical scattering is generally characterized by an x-rayphoton interacting with a whole atom so that the photon is scatteredwith substantially no change in internal energy to the scattering atom,nor to the x-ray photon. In other words, the wavelength of the scatteredphoton is generally similar to that of the incident photon prior tobeing scattered.

[0006] Various attempts have been made at reducing the incidence ofscatter in radiographic imaging. For instance, the use of anti-scattergrids in full field digital imaging as described above has been shown toreduce the amount of scatter displayed in a resultant image. However,use of these grids generally coincides with a need for significantincrease in tissue radiation dosages to generate images of desiredresolution, signal-to-noise ratio and/or contrast. As another example,various opaque shields have been utilized to enable collection andcalculation of an estimated scatter portion of the total beam intensity.However, use of these opaque shields has resulted in portions of theimage information being lost (unexposed, white areas on film). As a wayof avoiding these lost areas in the resultant images, one or moresubsequent images may be taken without the shields in an additionalexposure period(s), and the scatter calculations of the first image canthen be applied to the subsequent image(s). However, this obviouslyrequires additional exposure of the patient to radiation doses, andmeans or process to remove the opaque shields between exposures.

SUMMARY OF THE INVENTION

[0007] The present invention is directed to a radiographic imagingdevice and associated methodology for reducing effects related toscattered radiation or scatter. In this regard, the invention reducesdisplay artifact associated with scatter, resulting in improved contrast(for given imaging parameters) associated with features being imaged,for example, physiological features such as lesions, masses,calcifications and the like in the case of mammography. Such scattereffects are reduced according to the present invention by reducingdetected scatter (“scatter rejection”) and/or processing of detectorinformation to compensate for scatter effects potentially present insuch information (“scatter compensation”). This imaging improvement mayallow for various advantages including reduced patient dosages, improvedresolution and/or improved detection of low contrast features.

[0008] Various aspects of the invention may be advantageouslyimplemented in connection with a composite imaging system that obtainsimaging information for different portions of an area of interest atdifferent times and aggregates this imaging information to generatecomposite imaging information for the area of interest. One example ofsuch a composite imaging system is a slot scanner such as a digital slotscanner. Indeed, such slot scanners provide inherent advantages relatedto scatter rejection. Slot scanners utilize an imaging beam, such as afan-shaped beam, that is narrower (in at least one dimension) than thearea of interest to be imaged. To image the full area of interest, thenarrow beam is scanned across the area of interest over an exposure timeperiod. In the case of digital slot scanners, a composite image of thefull area of interest may be constructed from the imaging informationincrementally obtained during the exposure period.

[0009] In slot scanners, a narrow active detector area, e.g., closelycorresponding to the cross-section of the imaging beam at the detector,is preferably moved in concert with the scanning motion of the imagingbeam. In the case of digital detectors, this may be accomplished, forexample, by physically moving a narrow array of detector elements or byelectronically moving the active portion of a stationary full fieldarray in concert with the scanning motion. In any such case, significantscatter rejection is achieved because radiation outside of the plane(or, more precisely, narrow wedge) including the signal source and theactive detector area is not recorded as part of the image-forminginformation. It should be noted that pixels outside the area illuminatedby the primary beam, may be used according to the present invention todirectly record scatter-only information; such information may then beused to contribute (for example via interpolation) to the estimate ofthe scatter in the area adjacent that illuminated by the primary beam.

[0010] However, it has been recognized that scattered radiation,including single scatter radiation (associated with a single scatterevent/interaction) and multiple scatter radiation (associated with morethan one scatter event/interaction), associated with Rayleigh and/orCompton scattering, may still impinge on the active detector area,presenting a possibility of scatter related image degradation. Thepresent invention reduces such potential degradation through enhancedscatter rejection and/or scatter compensation.

[0011] In accordance with one aspect of the present invention, a methodand apparatus (“utility”) is provided for rejecting scatter in thecontext of a composite imaging system. In particular, the utilityinvolves operating a source to transmit a photonic imaging signal, suchas an x-ray beam, to an area of interest of a patient's body such as apatient's breast for mammographic procedures. A detector detects atleast portions of the imaging signal from the patient's body andprovides an output reflecting imaging information for portions of thearea of interest obtained at corresponding different times of anexposure period. A processor processes the detector output to providecomposite imaging information for the area of interest. The utilityfurther involves selectively blocking, on a spatially dependent basis,photons that are directed to a detector location by allowing passage ofphotons substantially on a first linear pathway between the source andthe detector location and blocking photons associated with a pathwaydisposed at an angle to the first pathway.

[0012] For example, a rejection assembly, including at least onerejection element, e.g., a blinder element, may extend into the secondpathway to block associated photons from reaching the detector location.The scatter rejection assembly is preferably positioned and/or operatedso as to minimize interference with the desired imaging signal whileeffectively reducing the effects of scatter. In this regard, a rejectionelement of the assembly is preferably separated from the first pathway.The height of the rejection elements is determined as a result of atrade-off optimization between several system parameters, includingair-gap and image spatial resolution. For anticipated digital slot scanmammography applications, it is anticipated that the height may be atleast about 5-10 mm.

[0013] In certain implementations, the rejection element is disposedwholly outside of an imaging area defined by the source and an activedetector area. Alternatively, the rejection element may be partiallyinside the imaging area. The rejection element preferably includes asurface that is generally aligned with the first pathway between thesource and the detector area. In the context of a slot scan system, therejection assembly may have one or more surfaces where the orientationof the surface is dependent on scan location. In one embodiment, therejection assembly is disposed at least partially outside of an imagingarea and moves across an arcuate path in concert with a scan to avoid orreduce shadowing or the appearance of grid lines (at least grid-lines ata non-zero angle with respect to the scan axis).

[0014] The scatter rejection assembly may include multiple scanrejection elements. These elements may be parallel and/or transverse toone another. In connection with a slot scan system where the scannedbeam has an elongate beam cross-sectional axis, the longitudinalrejection elements may be aligned with the beam axis, e.g., they may bedisposed adjacent to each longitudinal edge of the beam for movement inconcert with the beam scan. An associate method in accordance with thepresent invention involves using the scatter blocking assembly to blockscattered radiation substantially without blocking any portion of theprimary imaging signal, thus avoiding shadowing (or grid lines) andpatient dose increase. Additionally or alternatively, the longitudinalelements may be aligned with the scan axis (or at a non-zero angle tothe detector axis) to reduce the effects of scatter within the beamplane. In this regard, a grid of transverse rejection elements may beused in the slot scanning context. For example, a stationary arcuate(e.g., partial spherical) grid may be used. Alternatively, such a gridmay be mounted for movement in concert with the scan so as to reduce orsubstantially eliminate shadowing or grid lines at any non-zero anglewith respect to the scan axis. In accordance with another aspect of thepresent invention, a utility is provided for using exposure time imagingparameter measurements for processing the associated imaginginformation. Such imaging parameters include patient dependentparameters and imaging signal parameters that may vary on a proceduredependent basis such as signal intensity, spatial intensitydistribution, or associated dependent parameters (e.g., detectorvoltage) and tissue thickness, tissue composition, patient/detector airgap or other geometric parameters. Such imaging parameter measurementsmay be distinguished from artifact measurements such as measurements ofsignal noise, electronic noise or scatter. It has been found thatmeasurements of such parameters during an exposure time period canadvantageously be used in processing associated image information.

[0015] The associated utility involves: transmitting a photonic imagingsignal relative to the area of interest during an exposure period of aradiographic procedure; with the patient in position for the procedure,measuring at least a first imaging parameter value and a second imagingparameter value; detecting at least portions of the photonic imagingsignal from the area of interest and providing imaging information basedthereon; and operating a processor to process the imaging informationusing the first and second imaging parameter values. The imagingparameter values may be used for a variety of purposes includingoptimizing digital image quality, display, and digital diagnosticprocessing and scatter compensation.

[0016] The imaging parameter measurements may be imaging signalparameters measured at different times and/or different locations duringthe exposure period. For example, in the context of slot scan imaging,values representative of the intensity of radiation impinging on thedetector at different times/scan locations of the scan can be measured.In scanning or other contexts, values representative of the intensity ofradiation impinging on the detector at different locations can bemeasured. Such intensity measurements may be used to develop a spatialprofile of the imaging signal relative to one or more axes of theimaging area. This spatial profile may be used for a variety ofpurposes, including scatter compensation, as described below.

[0017] Alternatively or additionally, the imaging parameter measurementsmay yield geometric imaging parameter values obtained either within oroutside of the exposure period. For example, in the context ofmammography, the geometric variables that may be measured include breastthickness (e.g., based on the engaged position of a compression paddle),tissue composition/density (e.g., based on compression force or aresistance profile obtained relative to progressive engagement of acompression paddle), and air gap distance between the compressed breastand the detector. Again, such measurements may be used for a variety ofpurposes, including scatter compensation, as described below.

[0018] According to a further aspect of the present invention, a utilityis provided for measuring an imaging parameter value in an imagingsystem. The utility involves transmitting an imaging signal from animaging source to a detector, disposing a signal attenuator between thesource and detector and measuring an imaging parameter based on a signalportion transmitted through the signal attenuator. The signal attenuatormay have attenuation characteristics that vary on a spatially dependentbasis. For example, a signal attenuating material having a varyingthickness, such as an acrylic wedge, may be disposed in a desiredposition relative to the detector. Alternatively, a window of varyingopacity may be provided for this purpose. The varying attenuationfacilitates accurate measurements, including comparative measurements,under a variety of imaging conditions without saturation (e.g.,exceeding the dynamic range of analog-to-digital converters or otherprocessing components), thereby accommodating exposure periodmeasurements or other measurements where the system settings may bedetermined at least in part by factors other than the attributes orlimitations of the imaging parameter measurement components.

[0019] In one implementation, one or more such signal attenuators areused in the slot scanning context. For example, a longitudinal acrylicwedge having its major axis aligned with the scan axis may be used toallow for intensity measurements over the course of a scan. In thismanner, scan position dependent intensity variations, e.g., due tosource fluctuations or detector alignment errors, can be identified andused for any of various purposes. The major axis of such a wedge mayalternatively be oriented transverse to the scan axis to characterize abeam profile relative to that axis. The combination of scan axis andtransverse attenuators allows for, characterization of a two-dimensionalprofile of the imaging signal impinging on the organ/object to be imagedas may be desired.

[0020] According to another aspect of the present invention, an “airgap” between a patient and a detector surface is selected to reduce theeffects of scatter. The associated utility involves establishing amathematical model for modeling a magnitude of expected scatterdetection as a function of a distance between tissue being imaged and adetector surface, and using the mathematical model to set a distancebetween a tissue support structure and the detector surface. Thisdistance may be varied for particular imaging procedures based on otherparameters such as tissue thickness and composition, and is optimized asa function of several imaging system performance including image,contrast and spatial resolution. Alternatively, a fixed spacing betweenthe tissue support and detector surface may be optimized based on arange of expected imaging parameters. According to a related aspect ofthe invention, a mammographic imaging system has a spacing between abreast support and a detector surface of between about 1-mm and 40-mmand, more preferably, between about 2-mm and 30-mm.

[0021] According to a still further aspect of the present invention, autility is provided for compensating for the effects of the scatter,taking into consideration both single and multiple scatter. The utilityinvolves: transmitting radiation into a selected tissue region of apatient's body; detecting radiation from the selected tissue region ofthe patient's body, wherein the detected radiation includes a scatteredportion and a non-scattered portion; estimating first and second partsof the scattered portion, wherein the first part corresponds withphotonic energy passing through the selected tissue region with a singlescattering occurrence and the second part corresponds with photonicenergy passing through the selected tissue region with multiplescattering occurrences; obtaining radiographic image data in relation tothe detected radiation from the selected tissue region; and utilizingthe estimated first and second parts of the scatter portion to adjustthe radiographic image data. The step of estimating the first and secondparts of the scattered portion may be based on a mathematical model. Themodel may address Compton scatter effects as well as Rayleigh, scattereffects. Additionally, the model may utilize any of a variety of imagingparameters including, for example, a measured dimension of the tissueregion, a tissue composition or density of the selected tissue region, apower setting for the radiation source (with associated, known sourceemission spectra), a measured intensity of detected radiation, and anintensity profile of detected radiation or the like. These values may bedetermined for a particular radiographic procedure and, moreparticularly, may be measured at one or more times during an exposureperiod of the radiographic procedure.

[0022] According to another aspect of the present invention, scattercompensation is performed based on procedure specific imaging parametermeasurements. The associated utility involves: transmitting radiationinto the selected tissue region; detecting radiation from the selectedtissue region, where the detected radiation includes a scattered portionand a non-scattered portion; positioning a patient in a desired positionfor a radiographic procedure; with the patient positioned in the desiredposition, operating a parameter measurement device to measure, aprocedure specific value of a scatter related parameter and provide anoutput indicative thereof; obtaining radiographic image data in relationto detected radiation from the selected tissue region; and operating aprocessor to receive the output from the parameter measurement deviceand use the image specific value of the scatter related parameter toadjust the radiographic image data. For example, the scatter relatedparameter may be an imaging parameter such as one or more valuesrelating to the intensity of the transmitted or detected radiation, orthe parameter value may be a geometric parameter relating to the tissuethickness or composition or the spacing between the tissue and thedetector. The value may be measured during an exposure period of theradiographic procedure or otherwise with the patient positioned for theprocedure. In this manner, procedure specific values are measured forimproved scatter compensation.

[0023] According to yet another aspect of the present invention, autility is provided for performing scatter compensation in connectionwith a composite imaging system. As noted above, such composite imagingsystems may provide advantages related to scatter rejection. Scatterperformance can further be enhanced by providing scatter compensation inconnection with such a composite imaging system. Thus, an associatedutility involves: transmitting a photonic imaging signal relative to anarea of interest of a patient's body during an exposure period of aradiographic procedure; first operating a detector to detect portions ofthe imaging signal from the patient's body and provide a detector outputindicative thereof, where the detector output reflects imaginginformation for different portions of the area of interest obtained atcorresponding different times of the exposure period; establishingscatter compensation information for the radiographic procedure; andoperating a processor to process the detector output using the scattercompensation information so as to provide reduced scatter compositeimaging information of the area of interest. Preferably, the scattercompensation information is based on a mathematical model for scattercompensation that uses measurements of one or more scatter parameters.These scatter parameters are preferably measured on a procedure specificbasis and may include exposure time measurements. Additionally, thescatter compensation information may take into account single andmultiple scatter including both Compton and Rayleigh scatter effects.Parameters of the scatter estimation models may be optimized as a resultof extensive imaging of various test objects, including anthropomorphicphantoms, as well as clinical investigations.

[0024] According to another aspect of the present invention, a compositeimaging system such as a slot scanner is implemented with an imagingbeam having a scan axis dimension no greater than a correspondingdimension of an active detector area. Heretofore, such composite imagingsystems have generally employed over-collimated beams, i.e., beamshaving a cross-sectional scan axis dimension (width) greater than thedetector width, in order to avoid scan axis intensity modulation andassociated display artifact due to, for example, beam/detector alignmentvariations during the scan. Such alignment variations may be caused, forexample, by scan drive fluctuations. Such overcollimation may result ingreater patient dosages and increased scatter potential.

[0025] The present invention includes a utility utilizing a matched orundercollimated beam. The utility involves: a source system fortransmitting an imaging beam; a detector for detecting the beam from thearea of interest of the patient's body and providing an output thatincludes imaging information for different portions of the area ofinterest obtained at different times; and a processor for processing thedetector output to provide composite imaging information for the area ofinterest. The source system and detector are configured such that adimension of the beam at the detector surface, relative to a first axis,is no greater than the corresponding dimension of an active detectorarea. In this regard, the source system may include a collimator forproviding the desired beam width. This configuration may allow directdetector measurements of the scatter-only radiation, and either directestimation of the scatter in the primary field (via interpolation) orrefinement of the scatter estimation model parameters via fitting of theestimates to the actual measured values in the scatter-only field. Anydrive related intensity fluctuations can be reduced by appropriatemechanical elements for stiffening the drive, electrical elements forappropriate modification of a drive signal, and/or sensor elements toidentify exposure period intensity variations so as to -allow forcompensation during processing. In this manner, potential reductions inpatient dosage and scatter can be achieved.

BRIEF DESCRIPTION OF THE DRAWINGS

[0026] For a more complete understanding of the present invention andfurther advantages thereof, reference is now made to the followingDetailed Description taken in conjunction with the drawings, in which:

[0027]FIG. 1 is a perspective view of a slot-scan digital mammographysystem in accordance with the present invention;

[0028]FIG. 2 illustrates the scanning motion of the imaging system ofFIG. 1;

[0029]FIGS. 3A-3C are front cross-sectional views illustrating thescanning motion of the system of FIG. 1;

[0030]FIG. 4 is a perspective partial cutaway, partially exploded viewof a detector assembly of the system of FIG. 1;

[0031]FIGS. 5A and 5B are front and side cross-sectional views,respectively, of a detector assembly in accordance with the presentinvention illustrating certain scatter paths;

[0032]FIG. 6 is a front view showing the geometry of the source anddetector system in accordance with the present invention;

[0033]FIG. 7 is a front partially schematic view of an imaging system inaccordance with the present invention;

[0034]FIGS. 8-10 show various grid configurations for use in accordancewith the present invention;

[0035]FIG. 11A is a perspective partially cutaway view showing adetector housing and related components in accordance with the presentinvention;

[0036]FIG. 11B is a front view showing a detector assembly in accordancewith the present invention;

[0037]FIG. 11C is a side view showing an alternative attenuatorconfiguration in accordance with the present invention; and

[0038]FIG. 12 is a flow chart illustrating a process for reducing theeffects of scatter in accordance with the present invention.

DETAILED DESCRIPTION

[0039] In the following description, the invention is set forth in thecontext of a digital, x-ray slot-scan mammography system. Thisrepresents a particularly advantageous implementation of the inventionas mammography is a challenging medical imaging application, in terms ofboth required contrast and resolution, and digital slot-scanningprovides a number of advantages as discussed above related to reducedscatter, reduced dosage, enhanced resolution and digital displayenhancement. However, various aspects of the invention are more broadlyapplicable in other contexts including other medical imagingapplications.

[0040] The following discussion first provides a description of adigital slot-scan mammography system. Thereafter, various systemcomponents for reduced scatter generation and improved scatter rejectionare described, i.e., system adaptations to reduce the amount ofscattered photonic energy detected at a detector surface. Finally,various system adaptations and processing techniques are disclosed forobtaining enhanced image-related information including scattercompensation, e.g., adjusting detected values to compensate for scattereffects.

[0041] 1. Slot-Scan System

[0042]FIG. 1 illustrates one embodiment of a slot-scan system 10 whichmay be utilized in implementing the present invention. The operation ofsuch a system is described in detail in U.S. Pat. Nos. 5,917,881 and5,526,394, assigned to Fischer Imaging Corporation, which areincorporated herein by reference. The system 10 includes a monitoringstation 12 and imaging station 14 operatively interconnected thereto.The monitoring station 12 may be located adjacent to the imaging station14 as shown or at a remote location. The illustrated monitoring station12 includes a user interface 16 such as a keyboard, mouse or the like(e.g., for entering patient data), a user interface display 18 (e.g.,for displaying/selecting/accepting images), and a processor 20interconnected to the user interface 16, user interface display 18 andimaging station 14. Processor 20 is adapted to receive, process andstore image signals generated at the imaging station 14 and other imagerelated information, as will be discussed in detail below, and tocontrol various operations at the imaging station 14. The monitoringstation 12 may also include a radiopaque and optically transparentshield 22 for shielding medical personnel during observed patientimaging operations at the imaging station 14. Various processingfunctions are described below including image time or post acquisitionprocessing to compensate for scatter. It is noted that various aspectsof such processing may be performed by a separate processor such as at aremote image review station. Moreover, such processing may bedistributed over multiple platforms, for example, to enable a serverwith substantial computing resources to support multiple imageacquisition and/or review stations at one or more medical facilities. Inother cases, a stand-alone unit may provide all of the processingfunctionality. The illustrated imaging station 14 includes an x-rayradiation source 24, e.g., an x-ray tube with collimating and filteringoptics 26, for transmitting a narrow radiation beam 28. The radiationsource 24 may be disposed for controlled rotation about its longitudinalaxis, wherein the radiation beam 28 may be scanned across a selectedregion of a patient's body, in this case, a compressed breast. Adetector assembly, as described below, is housed within detector housing30.

[0043] The scanning motion of the beam 28 is illustrated in FIG. 2.Specifically, the beam 28 is shown in bold at one end of a scan motion,and in shadow at other points of the scan motion. FIG. 2 also shows adetector assembly 32 that is housed within the housing 30 of FIG. 1. Thedetector assembly 32 preferably includes an active detector area 38, forreceiving the beam 28 and providing an electrical output indicativethereof, that closely corresponds in size and shape to the cross-sectionof the beam 28 at the surface of the active detector area. As discussedabove, composite imaging systems such as the illustrated slot-scanningarrangement have inherent scatter rejection advantages, as much of thescattered radiation will not be detected. The present invention may beadvantageously implemented in connection with such systems.

[0044] In this regard, such an active detector area 38 is moved inconcert with the beam 28 through the scan motion so as to continuouslyreceive the beam 28. Such movement may be accomplished by electronicallymoving the active portion of a large stationary detector. Alternatively,and as shown, the detector may be physically moved through the scan.Specifically, the illustrated detector assembly 32 includes a shuttlestructure 34 that carries a narrow detector area 38 as will be describedin detail below. The shuttle structure 34 is associated with a headstructure 36 of the source 24 such that the shuttle structure 34 travelsthrough an arc or other path as the head structure 36 rotates about anaxis of rotation.

[0045] The shuttle structure 34 and head structure 36 may be driventhrough this scan in any appropriate way. Preferably, the structures 34and 36 are co-driven to facilitate concerted motion of the beam 28 andactive detector area 38. In this regard, the structures 34 and 36 may beinterconnected such that driving one of the structures 34 or 36 resultsin coordinated movement of the other. A telescoping interconnectionstructure or other arrangement may be provided to allow for varying thedistance between the structures 34 and 36 as may be desired. In theillustrated embodiment, a motor 40, such as a microstep motor, iscarried by the shuttle structure 34. A spindle is mounted on an outputshaft of the motor to roll on cam surface 42. Thus, driving the motor 40in response to a drive signal from a controller associated with theprocessor 20 (FIG. 1) causes the spindle to wind scan cables affixed tothe gantry . . . which, in turn, drives the shuttle structure 34 throughthe scan arc and rotates the head structure 36.

[0046] Referring again to FIG. 1, the imaging station 14 may be utilizedto obtain various imaging views in relation to the patient's immobilizedbreast defined, for example, by the beam axis at the centerpoint of ascan. To allow for such views, the illustrated station 14 includes animaging unit 44 rotatably mounted on a pedestal 46. Pedestal 46 enablesmotorized elevation of the imaging unit 44 along a vertical axis. Theimaging unit 44 includes the source 24, detector housing 30 and othercomponents mounted on a common support structure. This support structureis mounted for rotation about an axis extending between the imaging unit44 and the pedestal 46. Optionally, the imaging unit 44 may also betiltable relative to a vertical plane. Such motions may be motorized toautomatically position the imaging unit 44 in response to inputs enteredat the monitoring station 12 or elsewhere on the imaging station, and abrake mechanism may be provided to lock the unit 44 in the desiredlocation.

[0047]FIGS. 3A-3C illustrate an image acquisition process in relation toa patient's compressed breast 48. The patient's breast 48 is compressedfor the procedure to provide a more uniform thickness and to immobilizethe patient for improved imaging. In the illustrated embodiment, thepatient's breast 48 is captured between an upper surface 50 of thedetector housing 30 and compression paddle 52. The illustrated paddle 52is arcuate in shape such that a more uniform tissue thickness isprovided across the scan motion and is formed from radio-transmissivematerial. In addition, the paddle 52 is movable on tracks 54 relative toan imaging axis 56 (coinciding with a center axis of the scan motion).Such motion may be motorized. Moreover, a release may be provided toallow the paddle 52 to be quickly disengaged from the patient's breast48 as may be desired.

[0048]FIG. 4 shows details of the active detector 38 within the housing30. As noted above, the active detector 38 provides an electrical outputrepresentative of detected radiation. In this regard, x-ray radiationmay be directly detected or spectrally transformed prior to detection.The illustrated detector 38 includes a scintillator strip 58 foremitting and transmitting light in response to incident x-ray radiationand a detector strip 60 for detecting the light. The scintillator strip58 is optically coupled to detector strip 60 via a fiber optic plate 62.The illustrated detector strip 60 is formed from four detector surfaces64.

[0049] The detector strip 60 defines an array of detector elements.These elements may form, for example, a single line of elements ormultiple lines of elements aligned with a major axis 66 of the array,also referred to as the detector axis. For example, these elements maybe provided as part of a CCD array or other photoelectric transducerarray. In the case of a single line array, values proportional toaccumulated charge may be read out from each element at a frequencyselected to provide the desired resolution in relation to the detectormotion along the scan axis 68. In the case of multiple lines, a timedelay integration process may be implemented by shifting charge acrossthe array in synchronization with the scan motion for read out at thetrailing edge of the array (alternatively signals may be combineddigitally, “off-chip”). This enables charge to be integrated for thefull beam width without sacrificing resolution, thereby enablingpractical imaging with available x-ray sources. For example, suchsynchronization may be achieved by linking charge shifts to pulses of amicrostep motor driving the scan or the output of an encoder associatedwith the moving detector (or other moving component). It is apparentthat this slot-scan geometry will result in reduced scatter generationand substantial scatter rejection. However, certain enhancements arepossible to further improve such scatter avoidance as discussed below.

[0050] 2. Scatter Avoidance

[0051] The present invention provides high-resolution medical images(for example, digital mammographic images) by very significantlyreducing the amount of detected scatter that contributes to suchresultant images. The amount of scatter contributing to a radiographicimage typically depends primarily on the amount of scatter generatedduring image acquisition and the proportion of generated scatter that isdetected by the detection assembly. While the angular distribution ofscatters (due to the Compton and Rayleigh effects) may vary somewhatwith energy (e.g., intensity of the radiation), at first approximationthe Compton scatters may be considered to be uniformly distributed at agiven angle (that is, isotropically distributed). By contrast, theRayleigh scatter tends to be more “forward-peaked” due to themaintenance of the photonic energy associated with this phenomenon. Inboth cases, the proportion of scatters generated at a given site andrecorded in the resultant image may be characterized as depending on asolid angle defined by the detector assembly and the scatter location.This “solid angle” generally corresponds with a distance from thescattering center to the detector (the solid angle decreases with thesquare of the distance, hence the effectiveness of air-gaps in reducingscatter in x-ray images), and on the detector area.

[0052] The narrow detector used in slot-scanning, and therefore the muchreduced detector area, explains the effectiveness of this technology inrejecting scatter. For example, for a 6-cm compressed breast of 50/50composition (50% fibro-glandular and 50% adipose), thescatter-to-primary ratio (the quantity of interest in medical imagingthat determines the amount of contrast reduction introduced by scatter)is of the order of 15%, or comparable to what could be achieved usingonly a conventional grid. Moreover, the present invention, as analternative to solely using traditional grid designs, beneficially leadsto significant intrinsic dose savings. This means that the patient isexposed to less radiation (relative to conventional imaging processes)during image acquisition. Further, lower intrinsic doses allows forhigher (or “tighter”) spatial resolution with high signal-to-noiseratios. That is, lower impinging radiation intensity levels, incombination with the reduction in scatter contributing to the resultantimage, allow for sharper or clearer radiographic images at a givenpatient dose level.

[0053] For any scatter point in the breast being imaged, the smalldetector area leads to a small probability of detection. The mean a freepath of an x-ray in tissue is dependent upon the tissue attenuationcoefficient and the x-ray energy. Considering a mean energy of 20-kev,and a corresponding attenuation coefficient of the order of 0.5 per cm,the mean x-ray free-path is of the order of 2-cm. It is therefore seenthat as the breast thickness increases, the likelihood of a given x-raybeing scattered more than once increases. For both analysis and scatterreduction considerations, it is useful to separate scatters in twocategories, single scatters (having been scattered or deflected once)and multiple scatters. Due to the geometry of a slot-scanning system, itis clear that all single scatters detected originated within the breastvolume illuminated by the beam (at a given instant). It is also clearthat multiple scatters detected are very unlikely to have had their lastscatter event within the illuminated volume (this follows directly fromthe solid angle geometric consideration above, and from the fact thatmultiple scatters may be considered isotropic in first approximation).Accordingly, a large portion of potentially detected multiple scatterswill be directed to the detector at an angle with respect to the beamplane (the plane including the focal spot and the detector axis edge ofthe breast support). It is therefore apparent that two parallelcollimator slats on either side of the active detector area may beeffective in reducing the detection in multiple scatters, while havingno effect on the detected primary beam.

[0054] This is illustrated in FIGS. 5A and 5B where I₀ indicates anilluminating intensity of the imaging beam 28. A primary portion P ofthe beam 28 traverses the patient's breast 48 and impinges on the activedetector area 38. Pathways 70 illustrate single scatter; that is,photonic energy passing through the patient's breast with only a singlescatter occurrence. Paths 72 illustrate multiple scatter. It will beappreciated that, depending on the type of scatter involved, thepathways 70 and 72 may be defined by the motion of one or more than onephoton. As shown, a small portion of single scatter 70 will be directedto the active detector area 38 as a result of the slot-scan geometry.Most of the multiple scatter 72 will be directed to the active detectorarea 38 from outside of the illuminated portion of the patient's breast48.

[0055] In the illustrated embodiment, collimator slats 74 are providedat the leading and trailing edges (defined in relation to the scanmotion) of the active detector area 38. The collimator slats 74 extend aheight, h, above the detector's surface relative to the beam axis 75(center axis of the beam) and are formed from a material suitable forabsorbing or blocking x-ray radiation. These slats 74 extend into thepathway of certain scattered radiation 72 to reject the radiation. Inparticular, the slats 74 will reject radiation having an incidence angleless than a rejection threshold measured relative to the plane of thedetector surface. The rejection threshold depends on the height of theslats and the points on the detector under consideration. Generally,however, the effectiveness of the slats 74 in rejecting scatter willdepend on the height h in relation to the width, w, of the activedetector array. In this regard, the height, h, is preferably at least50% and more preferably at least 100% of the width w . In theillustrated embodiment, the width, w, is about 10 mm and the height h isabout 10 mm. It will be appreciated that the slats 74 are particularlyeffective in rejecting multiple scatter (originating mostly in largebreasts).

[0056] Another factor that affects the amount of scatter detected is thewidth of the x-ray beam. In particular, the amount of, scatter generatedis proportional to the volume of tissue illuminated by the x-ray beam28. Thus, scatter can be reduced by narrowing the beam 28, while keepingthe detected primary intensity unchanged for a given technique. However,conventional slot-scan systems have generally utilized overcollimatedbeams, i.e., beams having a width in excess of the active detectorwidth. Such overcollimation accommodates relative motion between thebeam 28 and active detector area that might otherwise result inmodulation of the detected intensity. Thus, for example, for an activedetector area that is 10 mm wide, a 12 mm beam width at the detectorwould allow for 1 mm of beam motion with respect to each side of thedetector. The penalty for such overcollimation would be a 20% doseincrease as compared to the dose that contributes to image formation aswell as an increase in scatter generation, as well as a 20% increase inscatter generation (and thus in S/P) as compared to a 10-mm wide beamfor the same amount of detected energy.

[0057] The present invention may utilize a matched or undercollimatedbeam as illustrated in FIG. 6. Specifically, the width of the beam atthe detector surface, w_(b), is no greater than the width of the activedetector area, w_(d). In this regard, narrowing the beam 28 reducesscatter generation but also, in the case of undercollimation, the beamwidth over which charge integration can occur. Accordingly, w_(b) ispreferably between about 0.7 w_(d) and 1.0 w_(d) and, more preferably,between about 0.9 w_(d) and 1.0 w_(d). Moreover, to ensure adequateexposure in the case of undercollimation, the x-ray source preferablyprovides and output of at least about 150 mA in the range of 26-32 kVpand, more preferably, an output of at least about 200 mA in the notedrange. The illustrated embodiment uses a rad 71s tube. The illustratedembodiment also utilizes a high efficiency scintillator for optimizedDQE, image quality, and high-resolution imaging. It has been found thatintensity modulation due to relative motion between the beam 28 and theactive detector surface 38 is most pronounced at the beginning and endof the scanning motion, i.e., associated with scan acceleration. It isnoted that such modulation is of little practical significance in thecontext of mammography as the ends of the scan motion are generallyoutside of the image area of interest. Nonetheless, such potential imageartifacts can be addressed by varying the drive signal to the scanmotor. In particular, such intensity modulation can be empiricallymodeled to identify relative motion transients. These transients can beused to develop a modified ramp up and ramp down drive signal to smoothsuch transitions. These developed drive, signals can be tested andimproved to reduce the associated intensity modulation. Such intensitymodulation can also be corrected in processing by using a referencesignal. Such a reference signal may be generated by profiling thedetected signal relative to the scan axis. For example, during a scan,intensity measurements at the anterior edge of the array (the edgefarthest forward from the patient's chest wall), which is generallyoutside of the illuminated tissue region, can be read out over thecourse of the scan and used as a reference intensity signal. Thus, basedon this intensity profile, the measured intensity can be scaled ornormalized on a line-by-line basis relative to the scan axis to reduceor substantially eliminate any intensity modulation due to relativemotion between the beam 28 and the active detector area 38.

[0058] However, it will be appreciated that the beam portion transmittedthrough the patient's tissue will be significantly attenuated inrelation to the beam portions detected at the anterior edge of thedetector array. The amount of such attenuation will depend on, interalia, the thickness and composition of the illuminated tissue. In orderto avoid saturation and provide useful intensity profile measurementsunder a range of conditions, an attenuator 76 (see also FIGS. 11A and11B) is disposed adjacent to the anterior edge of the active detectorarea 38. The attenuator 76 preferably has signal attenuationcharacteristics that vary relative to the detector axis 66 (FIG. 4) butdo not vary substantially relative to the scan axis 68. Such variationrelative to the detector axis 66 may be accomplished by providing avarying thickness of an x-ray attenuating material or providing a windowhaving a varying opacity. In the illustrated embodiment, the attenuatorcomprises an acrylic wedge oriented to provide a varying thicknessrelative to the detector axis 66. In this manner, it is anticipated thata line of detector elements at some location between the narrowest andwidest point of the wedge can be used to provide information fordeveloping an intensity profile without saturation under a range ofimaging conditions.

[0059] Scatter rejection can also be enhanced by controlling the size ofthe air gap between the illuminated tissue and the detector surface. Inparticular, as the Compton component of single scatter may be assumedrelatively uniformly distributed in angle, while the Rayleigh componentof single scatter is highly forward-peaked, it is apparent thatincreasing the distance between the breast support and the detector (theair gap) will significantly decrease the amount of detected Comptonscatters while decreasing by a smaller proportional amount the Rayleighscatters detected. In particular, increasing the air gap from 0centimeters as shown in the embodiments of FIGS. 1-4, to a fewcentimeters will very significantly decrease the amount of total scatterdetected. Moreover, based on a mathematical model as discussed below formodeling of scattered radiation, it is apparent that the air gap can betuned according to the measured compressed breast thickness to leveragethis effect. According to such model, the median plane through thecompressed breast thickness is placed at distance from the detectordetermined from the breast thickness, composition, and imagingparameters such as kVp.

[0060]FIG. 7 illustrates a system where such tuning can be accomplished.In the illustrated embodiment, the patient's breast 48 is not compressedagainst the upper surface 50 of the detector housing 30. Rather, thepatient's breast 48 is supported by a support plate 78 that is separatedfrom the upper detector surface 80 by a distance d. This distance may bevariable or may be fixed. In the case where this distance is fixed, themagnitude of this distance may be tuned relative to an average expectedtissue thickness. For example, in the case of mammography, d ispreferably between 5 and 30 and more preferably between 10 and 20-mm.The support plate 78 is preferably curved so that this distance issubstantially constant throughout a scan.

[0061] In the illustrated embodiment, this distance may be varied bymoving the support plate 78 and/or the detector housing 30 and/or thedetector within the housing along the beam axis 75 on tracks 54. In thisregard, moving the housing 30 has the advantage that the air gap can beselected after the tissue thickness has been measured without moving thepatient. However, moving the housing 30 (or the detector within thehousing) is structurally more difficult to implement and results in avarying beam width at the detector surface 80. Thus, in operation, thepatient's breast 48 may first be positioned on the lower support plate78. The compression paddles 52 may then be manually or automaticallylowered to compress the patient's breast 48. Once the breast 48 iscompressed in an imaging position, the thickness of the patient'scompressed breast 48 may be measured, e.g., based on markings associatedwith the paddle 52 or based on an output signal from a paddle drivemotor 80 or an encoder. This value can then be provided to the processor20 (FIG. 1) which determines an optimal air gap. The lower support plate78 and/or the detector housing 30 may then be manually or automaticallymoved to establish the desired air gap for optimized scatter rejection.The breast support plates and gantry elevation motions may besimultaneous and in opposite directions, so that as a result theabsolute patient position in three-dimensional space is unchanged.

[0062] Scatter rejection may further be enhanced by using a stationaryor reciprocating grid 82, e.g., disposed below the upper surface of thedetector housing or on the upper surface 80 of the active detector area.A variety of different grid configurations may be utilized in thisregard. FIG. 8 illustrates one such grid configuration where the grid82′ is composed of transverse slats 84, for example, aligned with thedetector and scan axes. Such a grid is analogous to a Bucky grid used infull field imaging applications. However, the illustrated grid 82′ iscontoured to generally match the shape of the housing 30 (FIG. 7). Inaddition, the major axis of each of the slats 84 may be aligned with thebeam axis 75 (FIG. 5). In the context of the illustrated slot-scanner,the slots may therefore have an orientation that is dependent on theirlocation relative to the scan axis and/or the detector axis. The variousslats 84 function in a manner analogous to the collimator slatsdiscussed above to block radiation not on the primary incidence pathwayi.e., the beam axis 75.

[0063]FIG. 9 shows an alternative grid configuration for slot-scanningapplications. The illustrated grid 86 includes only slats 88 that aregenerally aligned with the scan axis. As noted above, collimator slatsmay be provided in connection with the detector, reducing orsubstantially eliminating the need for slats aligned with the detectoraxis. Thus, the grid design of FIG. 9 reduces grid line and dosageissues in relation to the design of FIG. 8. Nonetheless, such stationaryor reciprocating grids result in grid lines in the image and/or requireincreased patient dosages. Moreover, such large grids may be difficultto reciprocate fast enough to provide the desired grid blurring atpreferred scan rates/exposures.

[0064]FIG. 10 shows a further alternative where slats 90 disposed at anon-zero angle, θ, with the scan axis are provided in connection withthe active detector area 38. For example, the grid of FIG. 10 may bedisposed on the detector assembly above the scintillator for movementwith the detector assembly. The illustrated angling of the slats 90blurs the slats 90 so that lines do not appear in the resulting image.Moreover, the blurring is accomplished via a unidirectional movement ofthe slats 90 together with the detector, as opposed to reciprocatingBucky-style movement, thereby eliminating the issues of source/griddrive synchronization, increased exposure period to allow sufficientblurring movement (though some dosage increase may be associated withgrid shadowing of the primary, signal), and potential grid drivemalfunctions. It will be appreciated that the various slats of Figuresabove with a scan axis component are useful for blocking scattertraveling in or near to the plane of the detector axis and beam axis.Grid slat angles, height, width, parameters are optimized as a functionof various system imaging parameters.

[0065] The present invention thus may include a number of components forreducing scatter generation and improving scatter rejection.Nonetheless, some amount of scatter may be detected. Accordingly, thepresent invention further reduces the effects of scatter through scattercompensation, e.g., post-detection processing to subtract scatter fromthe raw image data as discussed below.

[0066] 3. Scatter Compensation

[0067] A. Gross Scatter Estimates and Spatially Dependent Estimates.

[0068] Scatter compensation relates to estimating the amount of detectedscatter so that the image information from the detector can be adjustedto compensate for the effects of such scatter, e.g., via digitalsubtraction. This estimation may be based at least in part on scattermeasurements or may be based on a mathematical model involving variousscatter related parameters. With regard to scatter measurements, asnoted above, the imaging system of the present invention may utilize anundercollimated beam. That is, the beam width may be less than thedetector width. The extra “dark” detector area may be used to directlymeasure scatter in connection with certain detector read-out mechanisms.For example, a beam width of 8-mm and a detector width of 10-mm providesa 1-mm dark band at each detector edge which may be used for scatterdetection, provided that these detector areas are separately read out orotherwise distinguished from the integrated charge associated with theprimary signal. Such measurements may be interpolated to providespatially indexed scatter data. Such an interpolation process may not belimited to linear interpolation but, rather, may take into account thespatial dependence (e.g., isotropic or forward peaked) characteristicsof various scatter effects and grid performance.

[0069] With regard to mathematical modeling, substantial improvements inimage quality can be achieved through gross scatter estimatesindependent of specific beam/scatter spatial distribution information. Afurther improvement can be achieved by modeling a scatter distributionbased on reference measurements obtained during a calibration procedureor at some other test time. As discussed below, the present inventionprovides a still further improvement by taking certain measurementsconcurrent with image acquisition so that scatter compensation can bedetermined or scaled based on actual imaging conditions. The associatedcompensation models may be empirically and/or mathematically derived. Amodel is discussed below for addressing Rayleigh as well as Comptoneffects and single as well as multiple scatters.

[0070] It is clear that multiple scatters do not carry much informationwith respect to the origin of the first scattering event, andaccordingly the multiple scatter distribution will be fairly uniform,slow-varying in space, and could be corrected by subtracting a simpleconstant or slowly varying function from the detected intensitydistribution as discussed below (as, for example, a single constant orlinear model for each line (row) in the image, plus smoothnessconstraints). On the other hand, single scatters will carry somegeometric information. Indeed, the probabilistic properties of theCompton scatter distribution are accurately known (for a givenscattering material). That is, as a function of incoming x-ray energy,the differential tissue cross sections allow analytic descriptions ofthe angular and energetic distribution of the scattered x-rays. (Theanalytic function is given by the Klein-Nishina expression). TheRayleigh scattering distribution is known experimentally for mostmaterials of interest to within less than 10% error. Both of thesefunctions have been included in various software Monte-Carlo codes thatallow simulation of the scatter distribution to a very precise degree.Examples include the EGS4 code from Standford (Electron Gamma Showers,version 4, Standford Linear Accelerator Program), and TART2000.

[0071] It is clear from the above consideration that it isstraightforward to carry out analytically calculations of thescatter-to-primary ratio (S/P) for sheets of uniform materials. Thesecalculations can then be generalized to non-uniform materials, and fromthe estimates of scatter derived from the line-integrals of the x-rayattenuation coefficients, a correction method can be derived. Theseestimates can be checked, and the model parameters adjusted for optimalresults, in the following two complementary ways. First, experiments canbe carried out on the imaging system 10 (FIG. 1) with sheets of uniformdensity. By varying the distance from the sheets to the detector, anaccurate approximation to the solid angle function and average depth oforigin of scatters within the sheets may be determined. Accordingly, asimple correction model may be derived, and the parameters of this modelmay be further adjusted by carrying out Monte-Carlo experiments. Thesetwo complementary approaches will lead to a scatter-correction methodthat will allow subtraction of at least 70 % of the scatter in thedigital images of system 10.

[0072] The calculation of the line-integral through the object beingimaged is a key component for an accurate estimate of scatter. Thiscalculation in turn depends on taking the log of the ratio of detectedintensity to impinging (or illuminating) intensity. Both of thesequantities are measured in the system 10, and these measurement may befurther refined. Indeed, the beam profile may be measured during adaily-calibration using the read out of a detector during a calibrationscan or scans. This impinging beam can be further observed byintroducing at the beginning (and/or end) of the scan an attenuator, forexample, an acrylic wedge of known geometry and properties as describedabove, such that under all expected imaging techniques, a non-saturatedvector of data can be acquired representing the anterior-posterior beamintensity profile. Another attenuator, e.g., an acrylic wedge, locatedat the anterior edge of the field-of-view, and parallel to the scanaxis, can be used in similar fashion to track the intensity variationsduring the scan. Such wedges 92 are shown in FIGS. 3A-3C and FIG. 11A.

[0073]FIG. 11C shows an alternative wedge profile that may be utilizedfor one or more of the elements 76 and 92. As shown, the wedge 93includes an initial rectangular cross-section portion 95 and a taperedportion 97. In this manner, substantially constant attenuationcharacteristics are provided over the rectangular portion 95 andassociated detector elements can be readily integrated or read out toprovide averaged or composite information.

[0074] The information associated with the wedges 92 together with theoutput associated with anterior attenuator 76 allows for precisedetermination of the impinging illumination profile over the scan area.This information, in conjunction with the recorded digital data, allowsdirect log calculation (according to Beer's law) and determination ofthe line-integrals The x-ray variations introduced by the attenuators 76and 92 themselves can be easily characterized (e.g., empirically) andsubtracted from the digital image. Similarly, the high-resolutioncollimator blades (located at the x-ray source collimator) could bedesigned to include a semi-opaque material that would serve the samepurpose for high-resolution imaging.

[0075] Further, direct experimentation with blocks of uniform sheets, atany kVp, but with low mA to allow collection of non-saturated data, willprovide a further means of model development and improvement, byallowing comparison of x-ray profiles Oust outside a phantom, e.g., astandardized acrylic mass) with and without the uniform sheet materialin the beam. Other means of estimating the scatter distribution includethe use of beads and/or variable apertures (totally blocking the primaryand therefore allowing an estimate of the scatter intensity distributionin the bead shadow). Such approaches are effective in modeling alow-frequency scatter function that can be scaled to reflect specificimaging techniques (kVp, mA), and material thicknesses (as measured bythe system). Scatter correction then ensues by subtraction of the scaledlow-frequency scatter distribution.

[0076] It is important to recognize that in slot-scanning technology,due to the potential absence of a grid, the scatter distribution may nothave certain high-frequency components as observed with systems using atwo-dimensional detector and scatter-rejection grid. This is because ascatter rejection grid is most effective at removing scatter originatingat greater distance from the point under consideration in the image, assuch scatter will impinge on the detector with a larger angle andtherefore be more likely absorbed by the grid slats. Conversely, due tothe preferential rejection of multiple and Compton scatters in aslot-scanning system, the scatter frequency components may contain morehigh-frequency components than what would be observed on a grid-less, 2Darea detector. In summary, a significant part of the scatter inslot-scanning technology may be taken out with a relatively simple,low-frequency model. It should also be pointed out that part of thatslowly-varying scatter distribution is removed by the image presentationalgorithm conventionally executed as part of digital imaging processing,as this algorithm includes a type of unsharp masking, whereby thelow-frequency components of the image are subtracted or much attenuatedin the corrected image to improve high-frequency component contrast. Itshould also be recognized that use of time delay integration asdescribed above (TDI), when scatter-only measurements are directlyavailable, will provide further data regarding the spatial distributionof the scatter, and therefore can be leveraged to improve the efficacyof scatter-correction algorithms.

[0077] Further, the use of multi-views in mammography, tomosynthesis andthree-dimensional imaging is useful in refining the scatter distributionestimate by allowing precise determination of the location of specificscatterers (such as dense masses, micro-calcifications, etc) in a 3-Dvolume. In the multi-view case, a 3-D reconstruction of the linearattenuation coefficient leads to very precise, iterative scattercorrection, by allowing forward projection and calculation of the single(and by generalization, multiple of any given order) scatterdistribution (and correction by subsequent subtraction). Concurrentultrasound images may also be used in this regard for locatingscatterers and characterizing tissue composition. In any case, amathematical model for estimating scatter based on certain scatterrelated parameters may be developed for the imaging system as describedbelow.

[0078] B. Basis for Scatter Model.

[0079] Considering the patient's breast under compression, optionallyincluding the compression assembly, let I₀ denote the impinging x-rayintensity. Let P denote the exit intensity being detected (in theabsence of scatter), and let S be the detected scatter intensity,considering one detector element during an integration time Δt,illuminated by a pencil beam of x-ray covering exactly the area A of thepixel under consideration.

[0080] The goal of a first calculation is to determine the amount ofscatter intensity illuminating a point M′. At diagnostic x-ray energies,the total linear attenuation coefficient can be decomposed as follows:

μ(l,E)=[μ_(Compton)(l,E)+μ_(Rayleigh)(l,E)]+μ_(Photo−Electric)(l,E)  (1)

[0081] where the first term [in bracket] on the right-hand side of theequation is the total scatter component of the linear attenuationcoefficient, sum of the Compton and Rayleigh components, while the lastterm on the right-hand side is the photoelectric component,corresponding to the probability of total attenuation of the x-ray. Thelinear attenuation coefficient and its components are seen to bedependent upon the coordinate l along the path length (line-integral)being considered.

[0082] Accordingly the detected primary is given by: $\begin{matrix}{P = {\int_{Spectrum}^{\quad}{{I_{0}(E)}\exp \left\{ {- {\int_{pathL}^{\quad}{\left\lbrack {\mu \left( {l,E} \right)} \right\rbrack {l}}}} \right\} {E}}}} & (2)\end{matrix}$

[0083] and the total image signal recorded is given by the sum of theprimary P and (total) scatter S components:

I=P+S   (2-b)

[0084] In projection imaging, 3-D information is not known (regardingthe object composition) so that in a first approximation only theline-integral of the attenuation coefficient is measurable, as given bya simplification of equation (2) above:P = I₀  exp {−∫_(pathL)  [μ(l, E_(Eff))]l},

[0085] Where E_(eff) is the beam effective energy.

[0086] Knowing the total path-length L, as for example given by thecompression thickness indicated by the system, one obtains the followingestimate expression for the average linear attenuation coefficient atthe effective beam energy: $\begin{matrix}{\overset{\_}{\mu} = {{- \frac{1}{L}}{Log}\quad \frac{P}{I_{0}}}} & (3)\end{matrix}$

[0087] On a mammography system, the thickness indication L is estimatedby a combination of distance measurement and applied compression force.Other means of estimating the compressed thickness are available, suchas optical imaging, etc. Accordingly, a two-dimensional map ofcompressed breast thickness, L(x,y) may be obtained, and used on apoint-by-point basis in equation (3) and elsewhere in the scattercorrection model.

[0088] The attenuation coefficients are also known to be a function ofthe x-ray energy. Knowing the technique selected for a particularexamination, such as retained kVp and mAs, and based upon acharacterization of the x-ray tube spectrum, it is possible to model thebeam spectral shape at any point along an attenuation path of knowncharacteristics. In the absence of detailed knowledge about the 3-Dstructure of the object of interest, simple models provide approximateinformation that is sufficient for a first order estimation of thegenerated scatter.

[0089] A simple first order model follows. The tissues being imaged areassumed to be composed of water only. Based on the tabulated attenuationproperties of water, it is then possible to estimate the effectiveenergy Eeff such that: $\begin{matrix}{\overset{\_}{\mu} = {{{{- \frac{1}{L}}{Log}\quad \frac{P}{I_{0}}} \cong {{- \frac{1}{L}}{Log}\quad \frac{I}{I_{0}}}} = {\mu_{water}({Eeff})}}} & (4)\end{matrix}$

[0090] Note that due to the presence of detected scatter in theintensity measurement I (I=S+P), the resulting estimate F_(eff) will bebiased. The scatter correction process might therefore be madeiterative, where a second, improved estimate of E_(eff) ensues scattercorrection as described below.

[0091] A higher order model may also be employed, whereby the linearattenuation coefficient is decomposed onto two basis vectors. Knowledgeof the compressed thickness, emitted spectrum at selected kVp, anddetector response as a function of energy enables finding the respectivecoefficients of the two basis functions. More generally, by decomposingthe linear attenuation coefficient onto N (known) basis functionsμ_(i)(E): $\begin{matrix}{{{\mu \left( {l,E} \right)} = {\sum\limits_{i = 1}^{N}{{\alpha_{i}(l)} \times {\mu_{i}(E)}}}},} & (5)\end{matrix}$

[0092] where the α_(i)(l) are the decomposition coefficients, dependentupon the distance l along the path from the x-ray source to the pixelbeing considered, one obtains by substituting in (2) above:$\begin{matrix}{{P = {\int_{Spectrum}^{\quad}{{I_{0}(E)} \times {\exp \left\lbrack {- {\int_{L}^{\quad}{\sum\limits_{i = 1}^{N}{{\alpha_{i}(l)} \times {\mu_{i}(E)}{l}}}}} \right\rbrack}{E}}}},} & (6)\end{matrix}$

[0093] subject to (s.t.): $\begin{matrix}{{\sum\limits_{i = 1}^{i = N}{\alpha_{i}(l)}} = 1.} & (7)\end{matrix}$

[0094] By commuting the sum and integral signs in (6), and integratingboth sides of (7) over the path L: $\begin{matrix}{{P = {\int_{Spectrum}^{\quad}{{I_{0}(E)} \times {\exp \left\lbrack {- {\sum\limits_{i = 1}^{i = N}{{\mu_{i}(E)}{\int_{L}^{\quad}{{\alpha_{i}(l)}{l}}}}}} \right\rbrack}{E}}}},} & (8)\end{matrix}$

[0095] s.t.: $\begin{matrix}{{\int_{L}^{\quad}{\sum\limits_{i = 1}^{i = N}{{\alpha_{i}(l)}{l}}}} = L} & (9)\end{matrix}$

[0096] wherein (9) L represents L(x,y) for the pixel underconsideration.

[0097] Accordingly, by defining the N unknowns:A_(i) = ∫_(L)  α_(i)(l)l, i = 1, …  , N,

[0098] The following two equations are obtained: $\begin{matrix}{{P = {\int_{Spectrum}{{I_{0}(E)} \times {\exp \left\lbrack {- {\sum\limits_{i = 1}^{i = N}\quad {A_{i} \times {\mu_{i}(E)}}}} \right\rbrack}{E}}}},} & (10) \\{{\sum\limits_{i = 1}^{N}\quad A_{i}} = {L.}} & (11)\end{matrix}$

[0099] From linear system theory, it is well known that given Mequations (forming a non-singular system), M unknown may be determinedby system inversion. Accordingly the system of equations (10) and (11)supports the decomposition of the attenuation coefficients onto twobasis functions. Examples in general radiography would include bone andsoft-tissue; in mammography, these two basis functions may be chosen aswater and calcium, or as fibroglandular (dense) and adipose (fat)tissues.

[0100] Further, it is clear that addition of one or more measurements;such as obtained by changing the beam kVp, would lead to the followingsystem of M+1 equations: $\begin{matrix}{{P_{j} = {\int_{{Spectrum}_{j}}{{I_{0}(E)} \times {\exp \left\lbrack {- {\sum\limits_{i = 1}^{i = N}\quad {A_{i} \times {\mu_{i}(E)}}}} \right\rbrack}{E}}}},{j = 1},{\ldots \quad M}} & (12) \\{{\sum\limits_{i = 1}^{N}\quad A_{i}} = {L.}} & (13)\end{matrix}$

[0101] Accordingly, and for illustration in the context of mammography,with measurements of the line-integral attenuation at M=2 differentspectra (obtained for example either by changing the beam kVp and/orchanging the beam filtration), it is possible to decompose the tissuesonto three basis functions, such as fibroglandular tissue, adiposetissue, and calcium. Other basis functions choice are possible, and canbe tailored to the specific application.

[0102] In the following, a model, is proposed for the estimation ofdetected scatter, by decomposing scatters into single scatters andmultiple scatters.

[0103] C. Single Scatter Estimation.

[0104] The following analysis is conducted for the Compton components.The Rayleigh estimate would be obtained by substituting all quantities(such as differential scatter cross-sections) for Compton with thecorresponding quantity for Rayleigh.

[0105] The collision cross-section σ^(C) determines the probability thatan incident photon will undergo a Compton scatter. In a thin layer ofthickness dx, the probability of scattering is given by the fraction ofthe beam that is occluded by the scattering sites: $\begin{matrix}{{{{Proba\_ of}{\_ scatter}} = {{- \frac{\varphi}{\varphi}} = {n_{e} \times \sigma^{C} \times {x}}}},} & (14)\end{matrix}$

[0106] where φ is the photon fluence in photons per square centimeters,n_(e) is the electron density (cm⁻³), σ^(C) is the Compton collisioncross-section (cm²×g⁻¹), and μ_(C) is the Compton linear attenuationcoefficient. The electron density n_(e) is given by: $\begin{matrix}{{n_{e} = {N_{0} \times \rho \times \frac{Z}{A}}},} & (15)\end{matrix}$

[0107] with

[0108] N₀: Avogadro's number

[0109] ρ: material density

[0110] Λ: atomic weight;

[0111] Z: atomic number

[0112] The number of photons scattered per unit volume is then given by:$\begin{matrix}{\frac{N_{s}}{V} = {{- \frac{\varphi}{x}} = {\varphi \times n_{e}\sigma^{C}}}} & (16)\end{matrix}$

[0113] The differential cross-section$\left( \frac{\sigma^{C}}{\Omega} \right)_{\Psi}$

[0114] is defined such that dσ^(C) is the probability that an incidentphoton will be deflected into the elemental solid angle dΩ when passingthrough an attenuator containing one scattering per unit area, at angleψ. Accordingly: $\begin{matrix}{{\sigma^{C} = {2\pi {\int_{0}^{\pi}{\left( \frac{\sigma^{C}}{\Omega} \right)_{\Psi}\sin \quad \Psi \quad {\Psi}}}}},} & (17)\end{matrix}$

[0115] or differentiating (16) with respect to Ω: $\begin{matrix}{\left\lbrack \frac{d^{2}N_{s}}{{V}{\Omega}} \right\rbrack = {\varphi \times n_{e} \times {\left( \frac{\sigma^{C}}{\Omega} \right)_{\Psi}.}}} & (18)\end{matrix}$

[0116] The amount of energy d²E_(s) that is scattered into an elementarysolid angle dΩ from the elemental volume dV is: $\begin{matrix}{\left\lbrack \frac{d^{2}E_{s}}{{V}{\Omega}} \right\rbrack = {{{h}^{\prime} \times \varphi \times n_{e} \times \left( \frac{\sigma^{C}}{\Omega} \right)_{\Psi}} = {{h}_{0} \times \left( \frac{{h}^{\prime}}{{h}_{0}} \right) \times \varphi \times n_{e} \times {\left( \frac{\sigma^{C}}{\Omega} \right)_{\Psi}.}}}} & (19)\end{matrix}$

[0117] Accordingly, the amount of energy scattered by the object beingimaged into an annulus of radius dr located at distance r from M isgiven by: $\begin{matrix}{S_{C} = {\int_{l = 0}^{l = L}{\left\{ {{I^{\prime}(l)}{\exp \left\lbrack {{- \overset{\_}{\mu}}\quad \frac{\left( {L - l} \right)}{\cos \quad \Psi}} \right\rbrack}\left( \frac{d^{2}E_{s}}{{V}{\Omega}} \right){\Omega}} \right\} {{V(l)}}}}} & (20)\end{matrix}$

[0118] with:

dV(l)=A×dl.

[0119] where I′(l) represents the illuminating intensity at position lalong the line-integral, the term within the exponential is theattenuation from the elemental volume dV at position l along theline-integral to point M′, and the last term inside the curly bracketsrepresents the amount of energy scattered into an elementary solid angledΩ from the elemental volume dV as calculated above.

[0120] The solid angle dΩ is simply given by: $\frac{dA}{R^{2}},$

[0121] where dA=2π×r×dr and R²[L−l]²+r². Also, I′=I₀×exp(−{overscore(μ)}×l), so that by substituting one obtains: $\begin{matrix}{s = {\int_{t = 0}^{t = L}\left\{ {l_{0}{\exp \left( {{- \overset{\_}{\mu}} \times l} \right)}{\exp \left\lbrack {{- \overset{\_}{\mu}}\frac{\left( {L - l} \right)}{\cos \quad \Psi}} \right\rbrack} \times n_{e} \times h\quad _{0} \times \left. \quad{\left\lbrack {\left( \frac{{h}^{i}}{h}\quad \right)(\Psi) \times \left( \frac{\sigma^{C}}{\Omega} \right)_{\Psi}} \right\rbrack \left( \frac{2\pi \times r \times d\quad r \times {\cos (\Psi)}}{\left\lbrack {L - l} \right\rbrack^{2} + r^{2}} \right)} \right\} {{V(l)}}} \right.}} & (21)\end{matrix}$

[0122] where the term in the second square bracket inside the integralis the well-known Klein-Nishina function, that lists the differentialcross-sections as a function of incoming energy hν₀ and scatter angle ψ.The Klein-Nishina formula also gives the scattered photon energy, hν′,as a function of the incoming photon hν₀ and scattering angle ψ. Thetotal scatter radiation detected at a point M′ is then obtained by firstcalculating the relative contribution of the corresponding annulus topixel M′, and then integrating the result over all image rays. It isclear to those skilled in the art that several computation arrangementsand simplifications are possible to effectively carry out thecalculations.

[0123] D. Multiple Scatters Estimation.

[0124] It is clear that the larger the amount of attenuation, the largerthe amount of radiation scattered. However, due to self-attenuation, theamount of detected single scatters does not necessarily increase simplyas a function of object size. The total amount of scatter generatedwithin the object increases with object size. A model that has beenshown to work well in estimating the multiple scatter (as determined byMonte-Carlo calculations and verified by experimentation) is:Multiple ∝ [α + β × S_(C) + γ × S_(R)] × [⟨ɛ + δ × ∫_(L)μ(l)l⟩],

[0125] where α, β, γ, ε, and δ are constants, and the term in squarebrackets represent a weighted average of the line-integrals calculatedover a neighborhood of the pixel under consideration, and S_(C), S_(R),are the Compton and Rayleigh single scatter estimates respectively.

[0126] In practice, the multiple scatter component tends to vary fairlyslowly spatially, and the constants may be determined semi-empiricallyfrom laboratory experiments and/or Monte-Carlo simulations.

[0127] The total scatter estimate is given simply as the sum of thesingle scatter components, Compton and Raleigh, added to the estimate ofmultiple scatters. It is apparent that this model may utilize variousscatter parameters including geometric parameters, such as the air gapdistance and patient thickness and composition, and imaging parameterssuch as the spatial intensity profile of the beam. In one implementationof the system as shown in FIG. 7, at least some of these parameters aredetermined automatically at imaging time with the patient positioned foran imaging procedure and are recorded in memory 92 accessible via aprocessor 90 for performing scatter-related computations.

[0128] Specifically, outputs from the drive motor 80 and a resistancesensor 94 may be used to determine the compression force or pressureapplied to the patient or a profile of resistance as a function ofprogressive compression. For example, the sensor 94 may include a straingauge, pressure transducer or the like. The drive motor 80 and/orposition sensor 96 may provide an indication of the position of thepaddle 52. This in combination with similar outputs from a motor orposition sensor 98 associated with the support plate 78 can be used todetermine the tissue thickness. The tissue thickness distribution as afunction of pixel coordinates (x,y) may be estimated from appliedcompression force, paddle deflection model, optical or other means.Also, as discussed above, readings related to the various attenuatorscan be used to provide a spatial intensity profile for the image area.

[0129] The structure of FIG. 7, including patient support/compressionelements 52 and 78 that may be made independent of the motion of thesource/detector gantry, has advantages for various applications such asthree-dimensional imaging. For example, this structure allows formaintaining the patient in a fixed (compressed) position while rotatingthe gantry to acquire images from different projection orientations thatcan be processed to generate spatial images and to spatially localizeareas of interest, e.g., for minimally invasive sampling or treatment,or to guide surgeons.

[0130] An overall process 100 for reducing the effects of scatter inaccordance with the present invention is illustrated by the flow chartof FIG. 12. The process is initiated by positioning (102) the patientfor a mammographic procedure. This may involve positioning the patient'sbreast on a support plate in a desired position between the x-rayimaging source and the detector. A compression assembly may then beoperated (104) to compress the patient's breast. For example, acompression paddle on the side of the patient's breast opposite thesupport plate may be lowered to provide the desired compression. Thisaction may be manual or automatic. In the case of an automaticcompression assembly, this may involve activating the drive motor toapply a desired compression pressure or force.

[0131] To improve scatter compensation, the thickness and/or compositionof the patient's breast may be determined (106). The thickness may bedetermined and entered into the processor automatically or manually. Inthe case of manual determination, a scale provided in connection withthe movable compression paddle may be used to determine the compressedthickness. In the case of an automatic compression assembly, an outputfrom the compression motor or an encoder may be used to record thecompressed thickness. Similarly, the composition may be determinedmanually or automatically. In this regard, an experienced physician mayevaluate the tissue compensation and make an appropriate entry via auser interface. Alternatively, a resistance sensor associated with themovable compression paddle may be used to provide a measure of thecompression force applied or a profile of the resistance encounteredduring progressive compression of the patient's breast so as to providean indication of tissue composition. Alternatively or in complement,previous film and/or digital image of the patients may be processedautomatically to provide an estimate of tissue density, thereby allowingone additional term in the basis function decomposition (12). Thethickness and composition information is stored in memory for use by theprocessor in scatter compensation calculations.

[0132] The processor may then be operated to determine (108) an optimalair gap between the patient's breast and the detector surface. As notedabove, such optimization may be based on the at least the thickness ofthe patient's compressed breast. Once this optimal air gap isdetermined, the actual air gap of the system may be set (110) tominimize scattered detection. This may involve repositioning thedetector housing and/or the support plate, and/or the detector withinthe housing, and may be carried out with simultaneous gantry elevationcompensation so that the patient position remains unchanged during theprocess.

[0133] As noted above, various imaging geometry information may beoutput (112) to memory associated with the processor prior to, during orafter an image exposure. In any event, an imaging exposure and scan isinitiated (114) by activating the source and operating the scan motor todrive the detector and source through a scan motion. At the beginning ofthe scan motion, an initial beam axis profile relative to the detectoraxis may be obtained (116) by scanning across a detector accessattenuator and reading out the corresponding detector elements. Duringthe scan, a scan axis profile may be obtained (120) by reading out thedetector elements associated with a scan axis attenuator. Concurrentwith obtaining the scan axis profile in the illustrated process, theimage data is obtained (118) by reading out the detector elementsassociated with the image area. Finally, a final beam or detector axisprofile may be obtained (122) by reading out the detector elementsassociated with another detector axis attenuator disposed at the end ofthe scan. These various profile measurements may be utilized to developan overall beam intensity profile. For example, the detector axisprofile for a given point in the scan may be determined by interpolatingbetween the scan start and scan end profiles as measured using thedetector axis attenuators. This profile may then be scaled using theoutput from the scan axis attenuator such that profile information canbe obtained for each desired area or each point or pixel of the detectorarray over the course of a scan. Any available three-dimensionalinformation, e.g., based on stereo imaging or ultrasound measurements,may also be provided (124) to the processor at an appropriate time. Theassociated image data and imaging parameter information is output (126)to the memory associated with the processor for use in scattercompensation calculations.

[0134] The mathematical model discussed above is then used to perform(128) a single scatter estimate and perform (130) a multiple scatterestimate. These calculations may use the various outputs discussed aboveincluding beam intensity profile information, patient thickness andcomposition information and other geometric parameters such as air gapdistance. The single scatter estimate and a multiple scatter estimateare combined to determine a total scatter estimate which, in turn, isused to obtain (132) scatter correction values on a pixel-by-pixelbasis. The raw imaging information is then corrected (134) using thesevalues. For example, this correction may be implemented by a digitalsubtraction process. Profile information may also be used to correct(136) for any scan artifact, e.g., due to scan motion transients. Thecorrected or adjusted imaging data may then be displayed (138) on amonitor. The result is that the effects of scatter are substantiallyreduced allowing for improved contrast, improved resolution, reducedpatient dosages and improved diagnosis.

[0135] Those skilled in the art will now see that certain modificationscan be made to the assembly and methods herein disclosed with respect tothe illustrated embodiments, without departing from the spirit of thepresent invention. And while the invention has been described above withrespect to the preferred embodiments, it will be understood that theinvention is adaptable to numerous rearrangements, modifications, andalterations, and all such arrangements, modifications, and alterationsare intended to be within the scope of the appended claims.

What is claimed:
 1. A method for use in imaging an area of interest of apatient's body, comprising the steps of: transmitting a photonic imagingsignal relative to said area of interest of said patient's body duringan exposure period of a radiographic procedure; with said patient in animaging position for said radiographic procedure, measuring at least afirst imaging parameter value and a second imaging parameter value;detecting portions of said photonic imaging signal from said area ofinterest of said patient's body and providing imaging information basedthereon; and operating a processor to process said imaging informationusing said first and second imaging parameter values.
 2. A method as setforth in claim 1, wherein at least one of said first and second imagingparameter values is a value of a patient dependent parameter.
 3. Amethod as set forth in claim 2, wherein said patient dependent parameterrelates to a thickness of said area of interest of said patient's body.4. A method as set forth in claim 2, wherein said patient dependentparameter relates to a tissue composition of at least a portion of saidarea of interest of said patient's body.
 5. A method as set forth inclaim 2, wherein said patient dependent parameter relates to a distancebetween said area of interest of said patient's body and a detectorsurface for detecting said portions of said photonic imaging signal. 6.A method as set forth in claim 2, wherein said step of measuringcomprises receiving a signal indicating a position of a component of animaging system for imaging said area of interest of said patient's body.7. A method as set forth in claim 2, wherein said area of interest ofsaid patient's body comprises at least a portion of a patient's breastand said imaging geometry parameter relates to a compression applied tosaid patient's breast.
 8. A method as set forth in claim 1, wherein atleast one of said first and second imaging parameter values relates toan imaging signal parameter.
 9. A method as set forth in claim 8,wherein said at least one of said parameter values relates to anintensity of said detected photonic imaging signal.
 10. A method as setforth in claim 8, wherein said at least one of said parameter values isused to determine a signal intensity profile.
 11. A method as set forthin claim 1, wherein said first and second imaging parameter valuesrelate to an imaging signal parameter at first and second locations ofan imaging area.
 12. A method as set forth in claim 11, wherein saidstep of operating comprises using said first and second parameter valuesto establish an intensity profile of photonic imaging signal relative toa first axis of a detector for detecting said photonic imaging signal.13. A method as set forth in claim 11, wherein said first and secondimaging parameter values are used in establishing an intensity profileof said photonic imaging signal relative to first and second axis of adetector surface for detecting said photonic imaging signal.
 14. Amethod as set forth in claim 1, wherein said first and second imagingparameter values relate to an imaging signal parameter at first andsecond times of the said exposure period.
 15. A method as set forth inclaim 14, wherein said step of detecting comprises moving an activedetector area relative to said area of interest during said exposureperiod.
 16. A method as set forth in claim 1, wherein said step oftransmitting a photonic imaging signal comprises forming said signalinto a beam having an elongate cross-section and said first and secondimaging parameter values are used to establish an intensity profilerelative to an axis of said beam.
 17. A method as set forth in claim 1,wherein said photonic imaging signal is transmitted by a source anddetected by a detector, further comprising the step of disposing asignal attenuator between said source and said detector.
 18. A method asset forth in claim 17, wherein said detector comprises an activedetector area that moves relative to a scan axis during said exposureperiod and said attenuator comprises an elongate element having alongitudinal axis, and said step of disposing comprises orienting saidattenuator such that said longitudinal axis is parallel to such scanaxis.
 19. A method as set forth in claim 17, wherein said attenuatormoves with said active detector area.
 20. A method as set forth in claim17, wherein said detector comprises an active detector area that movesrelative to a scan axis during said exposure period and said attenuatorcomprises an elongate element having a longitudinal axis, and said stepof disposing comprises orienting said attenuator such that saidlongitudinal axis is transverse to such scan axis.
 21. A method as setforth in claim 17, further comprising the step of disposing a secondattenuator between said source and said detector.
 22. A method as setforth in claim 21, wherein said detector comprises an active detectorsurface that is movable across a range of scanning positions during saidexposure period, wherein said first attenuator is disposed at a firstend of said range of scan positions and said second attenuator isdisposed at a second end, opposite said first end, of said range of scanpositions.
 23. A method as set forth in claim 1, wherein said step ofoperating comprises using said first and second imaging parameter valuesto estimate an amount of scatter reflected in said imaging information.24. A method as set forth in claim 1, wherein said step of operatingcomprises providing a mathematical model for estimating an amount ofscatter reflected in said imaging information based on procedurespecific imaging geometry parameter values, and using said first andsecond imaging parameter values in said mathematical model to perform ascatter correction process relative to said imaging information.
 25. Amethod as set forth in claim 1, wherein said step of operating comprisesproviding a mathematical model for estimating an amount of scatterreflected in said imaging information based on procedure specific signalintensity parameter values, and using said first and second imagingparameter values in said mathematical model to perform a scattercorrection process relative to said imaging information.
 26. A method asset forth in claim 1, wherein said step of operating comprises usingsaid first and second imaging parameter values to establish an intensityprofile of said photonic imaging signal and estimating an amount ofscatter reflected in said imaging information based on said intensityprofile.
 27. A method as set forth in claim 1, wherein said step oftransmitting comprises forming said photonic imaging signal into anelongate beam having a longitudinal beam axis, said step of detectingcomprises moving an active detector surface relative to a scan axisduring said exposure period, and said step of operating comprisesestablishing a beam axis profile of said beam based on parameter valuesmeasured during said exposure period, establishing a scan axis intensityprofile of said beam based on imaging parameter values measured duringsaid exposure, and using said scan axis profile to scale said beam axisprofile so as to obtain an area intensity profile for an imaging area,and using said area intensity profile to adjust said imaginginformation.
 28. An apparatus for use in imaging an area of interest ofa patient's body, comprising: a source for transmitting a photonicimaging signal relative to said area of interest of said patient's bodyduring an exposure period of a radiographic procedure; a sensor systemfor measuring, with said patient in an imaging position for saidradiographic procedure, at least a first imaging parameter value and asecond imaging parameter value and providing an imaging parameter outputindicative thereof; a detector for detecting portions of said photonicimaging signal from said area of interest of said patient's body andproviding imaging information based thereon; and a processor for usingsaid imaging parameter output to process said imaging information. 29.An apparatus as set forth in claim 27, wherein said sensor systemcomprises a first sensor for sensing a patient dependent parametervalue. 30 An apparatus as set forth in claim 28, wherein said patientdependent parameter value relates to a thickness of said area ofinterest of said patient's body.
 31. An apparatus as set forth in claim29, wherein said patient dependent parameter value relates to a tissuecomposition of at least a portion of said area of interest of saidpatient's body.
 32. An apparatus as set forth in claim 29, wherein saidpatient dependent parameter value relates to a distance between saidarea of interest of said patient's body and a detector surface of saiddetector.
 33. An apparatus as set forth in claim 29, wherein said areaof interest of said patient's body comprises at least a portion of apatient's breast and said patient dependent parameter value relates to acompression applied to said patient's breast.
 34. An apparatus as setforth in claim 28, wherein said sensor system comprises a sensor forsensing an imaging signal parameter.
 35. An apparatus as set forth inclaim 34, wherein said imaging signal parameter relates to an intensityof said detected photonic imaging signal.
 36. An apparatus as set forthin claim 34, wherein said imaging signal parameter relates to a signalintensity profile of said detected photonic imaging signal.
 37. Anapparatus as set forth in claim 28, wherein said first and secondimaging parameter values relate to first and second locations of animaging area.
 38. An apparatus as set forth in claim, 37, wherein saidprocessor is operative for using said first and second parameter valuesto establish an intensity profile of the photonic imaging signalrelative to a first axis of said detector.
 39. An apparatus as set forthin claim 37, wherein said processor is operative for establishing anintensity profile of said photonic imaging signal relative to first andsecond axis of said detector.
 40. An apparatus as set forth in claim 28,wherein said first and second imaging parameter values relate to firstand second times of said exposure period.
 41. An apparatus as set forthin claim 28, wherein said detector includes an active detector that ismovable relative to said area of interest during said exposure period.42. An apparatus as set forth in claim 28, wherein said source comprisesa beam forming element for forming said signal into a beam having anelongate cross-section and said processor is operative for establishingan intensity profile relative to an axis of said beam.
 43. An apparatusas set forth in claim 28, further comprising a signal attenuatordisposed between said source and said detector.
 44. An apparatus as setforth in claim 43, wherein said signal attenuator has signal attenuationcharacteristics that vary in a known manner on a spatially dependentbasis relative to at least one axis of said attenuator.
 45. An apparatusas set forth in claim 43, wherein said detector comprises an activedetector area that moves relative to a scan axis during said exposureperiod and said attenuator comprises an elongate element having alongitudinal axis, wherein said attenuator is oriented such that saidlongitudinal axis is parallel to said scan axis.
 46. An apparatus as setforth in claim 43, wherein said detector comprises an active detectorarea that moves relative to a scan axis during said exposure period andsaid attenuator comprises an elongate element having a longitudinalaxis, wherein said attenuator is oriented such that said longitudinalaxis is transverse to said scan axis.
 47. An apparatus as set forth inclaim 43, wherein said attenuator is mounted to move together with saidactive detector area.
 48. An apparatus as set forth in claim 43, furthercomprising a second attenuator disposed between said source and saiddetector.
 49. An apparatus as set forth in claim 48, wherein saiddetector comprises an active detector area that is movable across arange of scan positions during said exposure period, wherein said firstattenuator is disposed at a first end of said range of scan positionsand said second attenuator is disposed at a second end opposite saidfirst end of said range of scan positions.
 50. An apparatus as set forthin claim 28, wherein said processor is operative for using said firstand second imaging parameter values to estimate an amount of scatterreflected in said imaging information.
 51. An apparatus as set forth inclaim 28, wherein said processor is operative for providing amathematical model for estimating an amount of scatter reflected in saidimaging information based on procedure specific imaging geometryparameter values and using said first and second imaging parametervalues and said mathematical model to perform a scatter correctionrelative to said imaging information.
 52. An apparatus as set forth inclaim 28, wherein said processor is operative for providing amathematical model for estimating an amount of scatter reflected in saidimaging information based on procedure specific signal intensityparameter values, and using said first and second imaging parametervalues and said mathematical model to perform a scatter correctionrelative to said imaging information.
 53. An apparatus as set forth inclaim 28, wherein said processor is operative for use in said first andsecond imaging parameter values to establish an intensity profile ofsaid photonic imaging signal and estimating an amount of scatterreflected in said imaging information based on said intensity profile.54. An apparatus as set forth in claim 28, wherein said source includesbeam forming structure for forming said photonic imaging signal into anelongate beam having a longitudinal beam axis, said detector comprisesan active detector area that is movable relative to a scan axis duringsaid exposure period, and said processor is operative for establishing abeam axis profile of said beam based on parameter values measured duringsaid exposure period, establishing a scan axis intensity profile of saidbeam based on imaging parameter values measured during said exposureperiod, using said scan axis profile to scale said beam axis profile soas to obtain an area intensity profile for an imaging area, and usingsaid area intensity profile to adjust said imaging information.
 55. Anapparatus as set forth in claim 28, wherein said sensor system comprisesa first sensor for sensing a first value of an imaging geometryparameter and a second sensor for sensing a value of an imaging signalparameter.
 56. A method for use in imaging an area of interest of apatient's body, comprising the steps of: transmitting a photonic imagingsignal relative to said area of interest of said patient's body duringan exposure period of a radiographic procedure; measuring a patientdependent variable for said radiographic procedure and providing ageometric variable output indicative thereof; detecting portions of saidphotonic imaging signal from said area of interest of said patient'sbody and providing imaging information based thereon; and using saidgeometric variable output to process said imaging information.
 57. Amethod as set forth in claim 56, wherein said step of measuringcomprises obtaining a value related to a thickness of said area ofinterest of said patient's body.
 58. A method as set forth in claim 56,wherein said step of measuring comprises obtaining a value related to atissue composition of at least a portion of said area of interest ofsaid patient's body.
 59. A method as set forth in claim 56, wherein saidstep of measuring comprises obtaining a value related to a distancebetween said area of interest and said patient's body and a detectorsurface for detecting portions of said photonic imaging signal.
 60. Amethod as set forth in claim 56, wherein said area of interest of saidpatient's body comprises at least a portion of a patient's breast andsaid step of measuring comprises obtaining a value related to acompression applied to said patient's breast.
 61. A method as set forthin claim 56, wherein said step of detecting comprises moving an activedetector area relative to said area of interest during said exposureperiod.
 62. A method as set forth in claim 56, wherein said step ofusing comprises performing a scatter correction to remove a scattercomponent of said imaging information.
 63. An apparatus for use inimaging an area of interest of a patient's body, comprising: a sourcefor transmitting a photonic imaging signal relative to said area ofinterest of said patient's body during an exposure period of aradiographic procedure; a sensor system for measuring a patientdependent variable for said radiographic procedure and providing apatient dependent variable output indicative thereof; a detector fordetecting portions of said photonic imaging signal from said area ofinterest of said patient?s body and providing imaging information basedthereon; and a processor for using said geometric variable output toprocess said imaging information.
 64. An apparatus as set forth in claim63, wherein said step of measuring comprises obtaining a value relatedto a thickness of said area of interest of said patient's body.
 65. Anapparatus as set forth in claim 63, wherein said step of measuringcomprises obtaining a value related to a tissue composition of said areaof interest of said patient's body.
 66. An apparatus as set forth inclaim 63, wherein said step of measuring comprises obtaining a valuerelated to a distance between said area of interest and said patient'sbody and a detector surface for detecting portions of said photonicimaging signal.
 67. An apparatus as set forth in claim 63, wherein saidarea of interest of said patient's body comprises at least a portion ofa patient's breast and said step of measuring comprises obtaining avalue related to a compression applied to said patient's breast.
 68. Anapparatus as set forth in claim 63, wherein said step of detectingcomprises moving an active detector area relative to said area ofinterest during said exposure period.
 69. An apparatus as set forth inclaim 63, wherein said step of using comprises performing a scattercorrection to remove a scatter component of said imaging information.70. An apparatus for use in imaging a selected tissue region of apatient's body, comprising: source means for transmitting radiationthrough a selected tissue region of a patient's body; receiving meansdisposed in opposing relation to said source means such that saidselected tissue region of said patient's body is positionabletherebetween, said receiving means comprising an array of detectorelements for accumulating electrical charge in relation to saidradiation; scanning means for scanning said receiving means by movingsaid array across said selected tissue region of said patient's body asradiation is transmitted through said selected region of said patient'sbody; profiling means for providing intensity information regardingradiation transmitted from said source means; and processing means forprocessing an output signal relating to said electrical charge and forcomposing a composite image indicative of said selected tissue region ofsaid patient's body.
 71. An apparatus as set forth in claim 70, whereinsaid source means comprises beam forming structure for forming saidtransmitted radiation into a beam having a dimension less than acorresponding dimension of said selected tissue region of said patient'sbody.
 72. An apparatus as set forth in claim 70, wherein said scanningmeans comprises mechanical means for moving said array across saidselected tissue region.
 73. An apparatus as set forth in claim 70,wherein said scanning means comprises circuitry for electronicallymoving said array across said selected tissue region of said patient'sbody.
 74. An apparatus as set forth in claim 70, wherein said profilingmeans comprises a portion of said array and detector elements.
 75. Anapparatus as set forth in claim 70, wherein said profiling meanscomprises detector elements for detecting radiation at differentlocations relative to said receiving means so as to profile andintensity of radiation instant on said receiving means relative to atleast one axis.
 76. An apparatus as set forth in claim 70, wherein saidprofiling means comprises detector elements for detecting radiation atdifferent locations relative to said receiving means so as to profileand intensity of radiation instant on said receiving means relative totwo axes.
 77. An apparatus as set forth in claim 70, wherein saidprofiling means comprises at least one radiation attenuator disposedbetween said source means and said receiving means, said radiationattenuator having radiation attenuation characteristics that vary in aknown manner on a spatially dependent basis relative to at least oneaxis of said attenuator.
 78. An apparatus as set forth in claim 70,wherein said processing means is operative for receiving an output fromsaid profiling means and using said output to adjust said compositeimage.
 79. An apparatus as set forth in claim 78, wherein saidprocessing means is operative for using said output to perform a scattercorrection relative to said composite image.
 80. A method of using amedical imaging system for imaging a tissue region of interest of apatient's body, comprising the steps of: emitting radiation from aradiation source toward an active array of detector elements; movingsaid active array of detector elements along a scan path substantiallyaligned with a support surface for placement of a portion of a patient'sbody including a selected tissue region to be imaged, said supportsurface being located between said radiation source and said array ofdetector elements; passing said radiation from said radiation source toa first intensity profiling structure during said emitting step;receiving said radiation from said radiation source using said array ofdetector elements and generated imaging information based thereon;generating radiation intensity profile data indicative of said radiationtransmitted to said profiling structure; and using said profile data toprocess said imaging information.
 81. A method as set forth in claim 80,wherein said step of emitting comprises forming said radiation into abeam having an elongate cross-section.
 82. A method as set forth inclaim 80, wherein said step of moving comprises mechanically moving adetector including said active array of detector elements.
 83. A methodas set forth in claim 80, wherein said step of passing comprisesattenuating said radiation on a spatially dependent basis.
 84. A methodas set forth in claim 80, wherein said step of generating comprisesestablishing a radiation intensity profile relative to an axis of saidscan path.
 85. A method as set forth in claim 80, wherein said step ofgenerating comprises establishing a radiation intensity profile relativeto an axis transverse to said scan path.
 86. A method as set forth inclaim 80, wherein said step of using comprises performing a scattercorrection relative to said imaging information based on said profiledata.
 87. A method for use in imaging a selected tissue region of apatient, comprising the steps of: transmitting photonic energy relativeto the selected tissue region of the patient during an exposure period;first operating at least a first detector portion of a detector toreceive at least a first energy portion of said photonic energy duringat least a first time portion of said exposure period, said first energyportion corresponding to a primary ray transmitted to said firstdetector portion free from scattering occurrences; and second operatingat least a second detector portion of the detector to receive at least asecond energy portion of said photonic energy during a second timeportion of said exposure period, said second energy portioncorresponding to scatter incident on said second detector portion afterone or more scattering occurrences.
 88. A method as set forth in claim87, wherein said first time portion and said second time portion atleast partially overlap.
 89. A method as set forth in claim 87, whereinsaid first detector portion of said detector comprises a first detectorsurface area and said second detector portion of said detector comprisesa second detector surface area separate from said first detector surfacearea.
 90. An apparatus for use in imaging an area of interest of apatient's body, comprising: a source system for transmitting an imagingbeam relative to said area of interest of said patient's body during anexposure period, said imaging beam having a first beam dimensionmeasured relative to a first axis; a detector for detecting portions ofsaid transmitted imaging signal that have interacted with said area ofinterest of said patient's body and providing a detector outputindicative thereof, said detector including an active detector areahaving a first detector dimension relative to said first axis, whereinsaid detector output includes imaging information for different portionsof said area of interest obtained at corresponding different times ofsaid exposure period; and a processor for processing said detectoroutput to provide composite imaging information of said area of interestof said patient's body; said source system and detector being configuredsuch that said first beam dimension at a location of incidence of saidimaging beam on said detector is substantially no greater than saidfirst detector dimension.
 91. An apparatus as set forth in claim 90,wherein said first beam dimension is less than said first detectordimension.
 92. An apparatus as set forth in claim 90, wherein saidsource system and said detector are configured such that said detectorincludes a first portion that receives said beam and a second portionoutside of a pathway of said beam, and said processor is operative forestimating scatter based on an output from said second portion of saiddetector.
 93. A method for use in imaging an area of interest of apatient's body, comprising the steps of: first transmitting an imagingsignal, having a first beam dimension measured relative to a first axis,through a first portion of said area of interest within said patient'sbody; first operating a detector including an active detector area todetect substantially an entirety of said first beam dimension andprovide first imaging information indicative thereof; first moving saidimaging signal relative to said first axis so as to transmit saidimaging signal through a second portion of said area of interest withinsaid patient's body; second moving said active detector area relative tosaid first axis; after said step of second moving, second operating saiddetector including said active area to detect substantially saidentirety of said first beam dimension and provide second imaginginformation indicative thereof; and using said first imaging informationand said second imaging information to provide a composite image of saidarea of interest.
 94. A method as set forth in claim 93, wherein saiddetector is further operative for detecting radiation outside of apathway of said imaging signal and providing a scatter output indicativethereof, said method further comprising using said scatter output inproviding said composite image.